The present invention relates to an ultrasonic visualization method and apparatus and is particularly useful in obtaining intravascular ultrasonic images. There is disclosed in our U.K. Patent Nos. 2,208,138; 2,212,267 and 2,246,632, systems for producing intravascular ultrasonic images and the present invention is concerned with providing various improvements to those systems.
Intravascular ultrasound is gaining acceptance as a clinical tool in the diagnosis and treatment of coronary artery disease. The technical problem of forming an image from an ultrasound transducer mounted at the tip of a catheter having a diameter of typically 1 mm has led to two distinct approaches. First there are catheters with a single rotating transducer. These are capable of producing high quality images but suffer from image distortion due to vibration of the tip and non-uniform rotation of the transducer, problems caused by the long, flexible rotating drive shaft. The other approach is to have a multi-element transducer. As there are no moving-parts this design does not suffer from the image distortions of the rotational approach, but has hitherto suffered from poor image quality due to an inability to perform the beam-forming satisfactorily. This invention is concerned with multi-element transducers, and concerns a new approach to the beam-forming problem, which will give high quality images.
Using multi-element transducers to form high-quality ultrasound images is well understood in external ultrasound. The usual method is to transmit on a group of elements, with the excitation signal to each element having an appropriate time delay so that the transmitted acoustic field is focused at a particular point in the field. After this excitation, the signal received on each element of the subgroup is appropriately delayed prior to summation to focus the receive field response a specific range from the transducer. The set of delays can be altered during the echo period so that the receive focus is appropriate for the instantaneous echo-range, a technique called dynamic focus.
This approach can generate images with narrow point spread functions and low side-lobe levels. However the hardware architectures it demands are not possible within the millimeter size constraints of catheter based ultrasound. The need to separately address several elements at once means either the provision of a separate cable to each element, or significant beam-forming circuitry incorporating the delays closely connected to the elements positioned at the transducer head. Neither of these approaches can be provided in the space permitted.
The limited space available does permit a limited number of signal cables and a simple multiplexer arrangement. Such an architecture permits the use of simple synthetic aperture techniques such as those disclosed in U.S. Pat. No. 4,917,097 (Proudian/Endosonics). In this technique transmit-receive signals from individual elements are obtained sequentially, stored in memory and then post-processed by a synthetic aperture algorithm which includes the time delays required to focus the data. This technique has the limitation that the dynamic range, defined as the signal ratio between the signal peak and the side-lobes outside the main beam, is too low for high-quality grey-scale imaging. (see FIG. 2)
A further disadvantage of this technique is that for reduction of side lobes the pitch of the elements has to approach one-half wavelength which is very expensive to achieve. At a frequency of 30 MHz, this means the elements have to be separated by twenty three microns, and one hundred and twenty eight are needed for a 1 mm diameter catheter. The micro-engineering of such a large number of elements and the associated interconnection problems would make such a design expensive. This approach has the additional disadvantages that the acoustic response of such small elements has low sensitivity, and the electrical impedance of the elements is high, creating difficulties in efficient electrical matching to the transmit and receive circuitry.
To overcome this O'Donnell has described an alternative synthetic aperture approach, in which groups of adjacent elements are connected together to create a group-element. (See O'Donnell and Thomas, L. J., "Efficient Synthetic Aperture Imaging from a Circular Aperture with possible application to Catheter-Based Imaging" IEEE Trans Ultrasonic, Vol. 39, No. 3, pages 366-380, 1992). Electronic rotation is performed by successively connecting in a new element on one end of the array, and disconnecting one on the other end. This has the advantage that the group-element has a lower electrical impedance and has greater acoustic sensitivity than a single element but it still requires a large number of elements to give good quality beam-profiles.
To improve the beam-forming capability of these synthetic aperture techniques O'Donnell has described the use of an optimal filtering technique. This uses an optimization technique to produce a set of filter coefficients which produce a beam profile with lower side-lobe levels than the standard synthetic aperture coefficients. The application of this technique is described mainly in relation to the group-element data acquisition but it can also be applied to single element synthetic aperture. In particular it can be used to correct for the poor beam profile obtained by catheters with lower numbers of elements, enabling these to produce higher dynamic range images. However, the use of group elements, or lower number of elements than optimum gives a beam profile with very high side lobes as shown in FIG. 2 and FIG. 3. The application of the optimization technique will improve the side-lobe levels, but there is a trade-off with main lobe width and noise immunity which this method must contend with. It is better to use a technique which produces better beam profiles to start with.
The ultrasonic visualization method and apparatus to which the present invention applies has a number of medical applications. These include those in which a so-called stent has been inserted into a patient's artery in order to restore or substantially restore the lumen of the artery to its original cross-section. Stents are well known devices for this purpose and once inserted into a patient's artery would remain there. There is, however, a risk that the wall of the artery at the location of the stent (typically a diseased area of that wall) will reform over and around the stent itself to again reduce the cross-section of the lumen of the artery. This phenomenon is known as "restenosis."
In using an ultrasound system of the kind to which the present invention relates, there is a problem in distinguishing echo signals from the stent itself and those from the surrounding tissue. More particularly, because the echo signals reflected from the stent are many orders of magnitude greater than those reflected from the surrounding tissue, the latter are swamped by the former with a result that the effective visualization of the surrounding tissue is not achieved.
As part of the design of the system according to the present invention filter coefficients must be calculated which have sufficiently low side-lobes to form accurate images from the cross-product signals. Using a fixed set of filter coefficients, the problematic situation shown in FIG. 18 is encountered. The strong reflector (e.g. a stent) is in the side-lobes of the current aperture and hence is suppressed relative to the signals of equal strength in the main lobe. However, when one is imaging tissue at the focal point, the spatial filtering may still be insufficient to adequately suppress the side-lobe signals. Hence the array gives an output from the side-lobe signal and correspondingly gives an image on the screen at the focal point. The clinician then has no means of knowing whether this signal is a side-lobe signal from the stent or a true tissue signal at the focal point.
As discussed earlier, although O'Donnell and Thomas have described an approach which reduces the size of the side-lobes, this may still not overcome the problem discussed immediately above. This is because either the side-lobes may still not be low enough or if they are, then there will be a wider main lobe, as discussed earlier. The latter will result in images of the stent which are "smeared" thus making it difficult to detect tissue which is close to the stent.